The present invention relates to magnetic resonance imaging systems, and particularly to the radio frequency coils used in such systems.
Magnetic resonance imaging (MRI) utilizes hydrogen nuclear spins of the water molecules in the human body, which are polarized by a strong, uniform, static magnetic field of the magnet (named B0—the main magnetic field in MRI physics). The magnetically polarized nuclear spins generate magnetic moments in the human body. The magnetic moments are aligned parallel to the direction of the main magnetic field, B0, in a steady state, and produce no useful information if they are not disturbed by any excitation.
The generation of a nuclear magnetic resonance signal (NMR) for MRI data acquisition is accomplished by applying a uniform radio frequency (RF) magnetic field (named the B1 field or the excitation field) orthogonal to B0. This RF field is centered on the Larmor frequency of protons in the B0 field and causes the magnet moments to mutate their alignment away from B0 by some predetermined angle. The B1 field is produced in the imaging region of interest by an RF transmit coil which is driven by a computer-controlled RF source and an RF power amplifier. During excitation, the nuclear spin system absorbs magnetic energy, and the magnetic moments precess around the direction of the main magnetic field. After excitation, the precessing magnetic moments will go through a process of free induction decay, releasing their absorbed energy and returning to the steady state. During free induction decay, NMR signals are detected by the use of a receive RF coil, which is placed in the vicinity of the excited volume of the human body. The NMR signal is the secondary electrical voltage (or current) in the receive RF coil that has been induced by the precessing magnetic moments of the excited protons in the human tissue. The receive RF coil can be either the transmit coil itself, or an independent receive-only RF coil. The NMR signal is used for producing images by using additional pulsed magnetic gradient fields, which are generated by gradient coils integrated inside the main magnet system. The gradient fields are used to spatially encode the signals and selectively excite a specific volume of the human body. There are usually three sets of gradient coils in a standard MRI system, which generate magnetic fields in the direction of the main magnetic field, varying linearly in the imaging volume.
In MRI, it is desirable for the excitation and reception to be spatially uniform in the imaging volume for better image uniformity. In a standard MRI system, the best excitation field homogeneity is usually obtained by using a whole-body volume RF coil for transmission. The whole-body transmit coil is the largest RF coil in the system. A large coil, however, produces lower signal-to-noise ratio (SNR or S/N) if it is also used for reception, mainly because of its greater distance from the signal-generating tissues being imaged. Since a high signal-to-noise ratio is most desirable in MRI, special-purpose receive coils are used for reception to enhance the S/N ratio from the volume of interest.
In practice, a well-designed specialty RF coil should have the following functional properties: high S/N ratio, good uniformity, high unloaded quality factor (Q) of the resonance circuit, and high ratio of the unloaded to loaded Q factors. In addition, the coil device should be mechanically designed to facilitate patient handling and comfort, and to provide a protective barrier between the patient and the RF electronics. A further way to increase the SNR is to replace the single specialty coil with an array of smaller coils and through the use of multiple receivers, add the signals together at the image construction stage. It is desirable to have a large RF transmit coil to create a uniform B1 field and to minimize as far as possible any distortion of the transmit field due to the presence of a receive coil. Of course the receive coil is also tuned to the Larmor frequency and without intervention this coil couples very strongly to the transmitter coil. It is almost universal practice to detune the receive coil during the transmit pulse by some means. A common method is to switch an inductor across one or more, of the numerous tuning capacitors that are included in the receive coil. This switching usually takes two forms: active and passive. Both are shown clearly by Edelstein in U.S. Pat. No. 4,620,155 and summarized in FIG. 1.
The active form uses a PIN diode pre-biased during transmit with a small dc current that forward biases the diode and creates sufficient stored charge in the junction so that it remains conducting during the whole cycle of a RF current waveform.
The passive alterative replaces the single PIN diode switch with a crossed pair of fast switching diodes. This crossed diode configuration avoids the need for dc bias because each half cycle of the current waveform is handled by one or other of the two diodes.
With active PIN diode coil detuning, the “on” resistance of the PIN diodes is very low and therefore the parallel resonant circuit formed by the coil tuning capacitor, the PIN diode and the detuning inductor has a high Q factor, and consequently high impedance. Active de-tuning has the disadvantage of the inconvenience and difficulty of feeding a dc current to all the diodes without either spoiling the RF characteristics of the receive coil, or distorting the B0 field by fields created by the dc bias current.
With passive coil de-tuning, there are no dc bias wires to route. Passive de-tuning has two main disadvantages. Firstly, the coil is not de-tuned until there is enough RF induced voltage to cause the switching diodes to conduct. However, the induced voltage can be several hundred volts and so the passive deturing is quite effective. Also it is not uncommon to include one active de-tuning circuit so that the coils can be de-tuned even in receive mode and the coil does not distort very weak RF pulses. Secondly, the conduction losses are higher in fast switching diodes than for PIN diodes, this result in a lower parallel or “blocking” impedance compared with the active version. Add to this the switching losses in the diodes and considerable heat is created when used in the most power demanding situations.